Injectable, pore-forming hydrogels for materials-based cell therapies

ABSTRACT

The invention provides compositions and methods to form pores in situ within hydrogels following hydrogel injection. Pores formed in situ via degradation of sacrificial porogens within the surrounding hydrogel facilitate recruitment or release of cells. Disclosed herein is a material that is not initially porous, but which becomes macroporous over time.

RELATED APPLICATIONS

This application is a continuation of U.S. patent application Ser. No. 13/877,572, filed on Nov. 19, 2013, which is a 35 U.S.C. § 371 national stage filing of International Application No. PCT/US2011/055174, filed on Oct. 6, 2011, which claims the benefit of priority under 35 U.S.C. § 119(e) to U.S. Provisional Application No. 61/390,594, filed on Oct. 6, 2010, which is incorporated herein by reference in its entirety.

STATEMENT AS TO FEDERALLY SPONSORED RESEARCH

This invention was made with Government support under NIH R37DE013033 awarded by the National Institutes of Health and MRSEC DMR-0820484 awarded by the National Science Foundation. The Government has certain rights in the invention.

INCORPORATION-BY-REFERENCE OF SEQUENCE LISTING

The contents of the text file named “29297-082N01US_ST25.txt”, which was created on Jul. 10, 2013 and is 1 KB in size, is hereby incorporated by reference in their entirety.

FIELD OF THE INVENTION

This invention relates to biocompatible hydrogel compositions.

BACKGROUND OF THE INVENTION

Over the recent decades, biocompatible polymers have been used to form scaffolds that act as carriers for cell transplantation, or to recruit host cell populations into the device.

SUMMARY OF THE INVENTION

The invention provides compositions and methods to form porous hydrogels. For example, pores are formed in situ within hydrogels following hydrogel injection into a subject. Pores that are formed in situ via degradation of sacrificial porogens within the surrounding hydrogel (bulk hydrogel) facilitate recruitment or release of cells. For example, the resulting pore is within 5% of the size of the initial porogen.

Disclosed herein is a material that is not initially porous, but which becomes macroporous over time resident in the body of a recipient animal such as a mammalian subject. These compositions are associated with significant advantages over earlier scaffold compositions. The hydrogels described herein are well-suited to initially protect transplanted cells from host inflammatory responses, and then release transplanted cells after inflammation has subsided (e.g., after 12 hours, or 1, 3, 5, 7, or 10 days or more post-transplantation, i.e. resident in the body of the recipient). The hydrogels described herein also double as a surgical bulking agent, further minimizing inflammation in the host, and then later releasing cells.

Accordingly, the invention provides a composition comprising a first hydrogel and a second hydrogel, wherein the first hydrogel degrades at least 10% faster (e.g., at least 15%, at least 20%, at least 25%, at least 30%, at least 35%, at least 40%, at least 45%, or at least 50% faster) than the second hydrogel and wherein either the first hydrogel or the second hydrogel (or both) comprises an isolated cell. Preferably, the first hydrogel comprises a porogen that degrades leaving a pore in its place. For example, the first hydrogel is a porogen and the resulting pore after degradation insitu is within 25% of the size of the initial porogen, e.g., within 20%, within 15%, or within 10% of the size of the initial porogen. Preferably, the resulting pore is within 5% of the size of the initial porogen. The first hydrogel degrades more rapidly than the second hydrogel, because the first hydrogel is more soluble in water (comprises a lower solubility index). Alternatively, the first hydrogel degrades more rapidly because it is cross-linked to protease-mediated degradation motifs as described in U.S. Ser. No. 10/980,989 to Zilla, incorporated herein by reference).

The molecular mass of the polymers used to form the first hydrogel composition (a porogen) are approximately 50 kilodaltons (kDa), and the molecular mass of the polymers used to form the second hydrogel composition (bulk) comprises approximately 250 kDa. A shorter polymer (e.g. that of a porogen) degrades more quickly compared to that of a longer polymer (e.g., that of the bulk composition). Alternatively, a composition is modified to render it more hydrolytically degradable by virtue of the presence of sugar groups (e.g., approximately 3-10% sugar of an alginate composition). In another example, the porogen hydrogel is more enzymatically degradable compared to the bulk hydrogel. The composite (first and second hydrogel) composition is permeable to bodily fluids, e.g., such as enzyme which gain access to the composition to degrade the porogen hydrogel. In some cases, the second hydrogel is cross-linked around the first hydrogel, i.e., the porogens (first hydrogel) are completely physically entrapped in the bulk (second) hydrogel.

Cells or bioactive factors (e.g., growth factors such as granulocyte/macrophage colony stimulating factor (GM-CSF), vascular endothelial growth factor (VEGF), condensed oligonucleotides, e.g., CpG, or plasmid DNA) are optionally encapsulated either into the porogen phase, bulk hydrogel phase, or into both phases. The porogens degrade in situ over a time-course pre-determined by the user. Upon degradation of the porogens, cells are released from or may migrate into the material. However, because they initially lack pores, pore-forming hydrogels are useful to provide mechanical support immediately after formation. Suitable bioactive factors include vascular endothelial growth factor (e.g., VEGFA; GenBank Accession Number: (aa) AAA35789.1 (GI:181971), (na) NM_001171630.1 (GI:284172472), incorporated herein by reference), acidic fibroblast growth factor (aFGF, Genbank Accession Number: (aa) AAB29057.2 (GI:13236891), (na) NM_000800.3 (GI:222144219), incorporated herein by reference), basic fibroblast growth factor (bFGF; GenBank Accession Number: (aa) AAB21432.2 (GI:8250666), (na) A32848.1 (GI:23957592), incorporated herein by reference), placenta growth factor (PIGF or PLGF; GenBank Accession Number: (aa) AAH07789.1 (GI:14043631), (na) NM_002632.4 (GI:56676307), incorporated herein by reference), leptin (Genbank Accession Number: (aa) CBI71013.1 (GI:285310289), (na) NM_000230.2 (GI:169790920), incorporated herein by reference), hematopoietic growth factor (e.g., HGF, Genbank Accession Number: (aa) AAA64297.1 (GI:337938), (na) NM_000601.4 (GI:58533168), incorporated herein by reference), VEGF receptor-1 (VEGFR-1, Genbank Accession Number: (aa) NP_002010.2 (GI:156104876), incorporated herein by reference), VEGFR-2 (Genbank Accession Number: (aa) AAC16450.1 (GI:3132833), (na) EU826563.1 (GI:194318421), incorporated herein by reference), transforming growth factor-β (TGF-β, Genbank Accession Number: (aa) AAA36738.1 (GI:339564), (na) NM_000660.4 (GI:260655621), incorporated herein by reference), bone morphogenetic protein (e.g., BMP-4, Genbank Accession Number: (aa) NP_570912.2 (GI:157276597), (na) NM_001202.3 (GI:157276592), incorporated herein by reference), insulin-like growth factor (IGF-1, Genbank Accession Number: (aa) CAA01954.1 (GI:1247519), (na) NM_001111283.1 (GI:163659898), incorporated herein by reference), fibroblast growth factor-2 (FGF-2), platelet-derived growth factor (PDGF; GenBank Accession Number: (aa) AAA60552.1 (GI:338209), (na) NM_033023.4 (GI:197333759), incorporated herein by reference), epidermal growth factor (EGF, Genbank Accession Number: (aa) AAH93731.1 (GI:62740195), incorporated herein by reference), transforming growth factor-α (TGF-α, Genbank Accession Number: (na) NM_003236.2 (GI:153791671), incorporated herein by reference), nerve growth factor (NGF, Genbank Accession Number: (aa) AAH32517.2 (GI:34192369), (na) NM_002506.2 (GI:70995318), incorporated herein by reference), brain-derived neurotrophic factor (BDNF, Genbank Accession Number: (aa) CAA62632.1 (GI:987872), (na) NM_170731.4 (GI:219842281), incorporated herein by reference), neurotrophin-3 (NT-3, Genbank Accession Number: (aa) NP_001096124.1 (GI:156630995), (na) NM_001102654.1 (GI:156630994), incorporated herein by reference), ciliary neurotrophic factor (CNTF, Genbank Accession Number: (aa) AAB31818.1 (GI:633830), (na) NM_000614.3 (GI:209574322), incorporated herein by reference), and glial cell line-derived neurotrophic factor (GDNF, Genbank Accession Number: (aa) CAG46721.1 (GI:49456801), (na) NM_000514.3 (GI:299473777), incorporated herein by reference). Other suitable factors include anti-VEGF antibody, anti-aFGF antibody, anti-bFGF antibody, anti-PIGF antibody, anti-leptin antibody, anti-HGF antibody, anti-VEGFR-1 antibody, anti-VEGFR-2 antibody, batimastat (BB-94), marimastat (BB-2516), thalidomide, O-(chloroacetylcarbamoyl)-fumagillol (TNP-470), carboxyamidotriazole (CAI), mitoxantrone, doxorubicin, SU5416, anti-TGF-β antibody, anti-BMP antibody, anti-IGF-1 antibody, anti-FGF-2 antibody, anti-PDGF antibody, anti-EGF antibody, anti-TGF-α antibody, and anti-VEGF antibody. Other bioactive factors suitable for encapsulation either into the porogen phase, bulk hydrogel phase, or into both phases include FMS-like tyrosine kinase 3 ligand (Flt3 ligand; Genbank Accession Number: (aa) AAI44040 (GI:219519004), (na) NM_004119 (GI: GI:121114303), incorporated herein by reference), anti-flt3 ligand, hepatocyte growth factor (Genbank Accession Number: (aa) AAB20169 (GI:237997), incorporated herein by reference), and stromal derived factor 1 (SDF-1).

Alternatively, an adenovirus is optionally encapsulated either into the porogen phase, bulk hydrogel phase, or into both phases. For example, the adenovirus encodes runt-related transcription factor (e.g., Runx2; Genbank Accession Number: (aa) CAI13532 (GI:55959066), (na) NM_001024630 (GI:226442782), incorporated herein by reference), a key transcription factor associated with osteoblast differentiation. In another aspect, the adenovirus encodes MyoD (Genbank Accession Number: (aa) CAA40000 (GI:34862), (na) NM_002478 (GI:77695919), incorporated herein by reference), a protein with a key role in regulating muscle differentiation. Alternatively, the adenovirus encodes bone morphogenetic protein, e.g., BMP-2 (Genbank Accession Number: (aa) AF040249_1 (GI:6649952), (na) NM_001200 (GI:80861484), incorporated herein by reference) or BMP-4 (Genbank Accession Number: (aa) NP_570912.2 (GI:157276597), (na) NM_001202.3 (GI:157276592), incorporated herein by reference). BMP-2 is involved in, inter alia, bone repair, while BMP-4 is involved in the repair of cardiac tissue. In one aspect, an adenovirus that encodes Runx and an adenovirus that encodes BMP-2 are encapsulated into the hydrogel.

Cells suitable for being encapsulated either into the porogen phase, bulk hydrogel phase, or into both phases include mesenchymal stem cells, myoblasts, vascular progenitor cells (e.g., an outgrowth endothelial cell), differentiated cells derived from embryonic stem cells or induced pluripotent stem cells, induced pluripotent cells, or cells that were directly reprogrammed from a fibroblast to a differentiated state.

In some examples, the porogen composition comprises cells, and in other examples, the bulk composition comprises cells. If cells are present in the composition (e.g., having been seeded during fabrication), the cells are deployed out of the composition after administration into a mammalian subject. Alternatively, the composition does not comprise cells; however, upon administration into tissues of a mammalian subject (e.g., implantation into a human patient), cells are recruited into the composition. The mammal can be any mammal, e.g., a human, a primate, a mouse, a rat, a dog, a cat, a horse, as well as livestock or animals grown for food consumption, e.g., cattle, sheep, pigs, chickens, and goats. In a preferred embodiment, the mammal is a human. Alternatively, the subject can be a non-mammalian animal such as xenopus, salamander, or newt.

The invention provides methods of deploying cells from a scaffold into tissues of a mammalian subject, comprising administering to a subject a composition comprising a first hydrogel and a second hydrogel, wherein the first hydrogel degrades at least 10% faster than the second hydrogel, and wherein the composition lacks pores at the time of administration, and wherein the composition comprises pores following residence in said subject, and wherein the first hydrogel or the second hydrogel comprises an isolated cell.

A methods of recruiting cells into a scaffold in vivo is carried out by administering to a subject a composition comprising a first hydrogel and a second hydrogel, wherein the first hydrogel degrades at least 10% faster than the second hydrogel and wherein the composition lacks pores at the time of administration, but comprises pores following residence in the subject. For example, pores are created due to the relative degradability or solubility of a first hydrogel composition compared to a second hydrogel composition, e.g., a porogen composition compared to a bulk composition.

Porosity influences recruitment and/or egress of the cells from the composition. Pores are nanoporous, microporous, or macroporous. For example, the diameter of nanopores is less than about 10 nm. Micropores are in the range of about 100 nm to about 20 μm in diameter. Macropores are greater than about 20 μm (e.g., greater than about 100 μm or greater than about 400 μm). Exemplary macropore sizes include 50 μm, 100 μm, 150 μm, 200 μm, 250 μm, 300 μm, 350 μm, 400 μm, 450 μm, 500 μm, 550 μm, and 600 μm. Macropores are those of a size that permit a eukaryotic cell to traverse into or out of the composition. In one example, a macroporous composition has pores of about 400 μm to 500 μm in diameter. The preferred pore size depends on the application. For example, for cell deployment and cell release, the preferred pore diameter is greater than 50 μm.

The size of the porogen is related to the size of the overall composite material. For example, for the material to stay intact, the porogen diameter is <10% of the smallest dimension of the overall composite. The density of porogens is between 10-80 percent of the overall volume of the composite composition. For example, the density of porogen is between 15% and 75%, between 20% and 70%, between 25% and 65%, between 30% and 60%, or between 35% and 55% of the overall volume. Preferably, the density of porogens is at least 50% of the overall volume to achieve optimal cell recruitment to the hydrogel or cell release from the hydrogel.

The hydrogel has an elastic modulus of between about 10 to about 1,000,000 Pascals (e.g., from about 10 to about 100,000 Pa, from about 10 to about 150,000 Pa, from about 10 to about 200,000 Pa, from about 10 to about 300,000 Pa, from about 10 to about 400,000 Pa, from about 10 to about 500,000 Pa, from about 10 to about 600,000 Pa, from about 10 to about 700,000 Pa, from about 10 to about 800,000 Pa, or from about 10 to about 900,000 Pa). Preferably, the slowly-degrading hydrogel comprises an elastic modulus of about 20 kilo Pa to 60 kPa, e.g., 25 kPa to 55 kPa, 30 kPa to 50 kPa, or 35 kPa to 45 kPa. The rapidly-degrading hydrogel comprises an elastic modulus of at least 40 kPa initially in order to maintain integrity during encapsulation prior to degradation.

Preferably, the slowly-degrading hydrogel (i.e., the second hydrogel or “bulk”) comprises high molecular weight peptides with an amino acid sequence of RGD which mimic cell adhesion proteins. Alternatively, the slowly-degrading hydrogel comprises a different adhesive peptide amino acid motif such as PHSRN (SEQ ID NO: 1) or DGEA (SEQ ID NO: 2). For example, the slowly-degrading hydrogels are preferably modified with 2-10 RGD peptides/polymer (e.g., alginate polymer).

By “hydrogel” is meant a composition comprising polymer chains that are hydrophilic. Exemplary hydrogels are comprised of materials that are compatible with cell encapsulation such as alginate, polyethylene glycol (PEG), PEG-acrylate, agarose, and synthetic protein (e.g., collagen or engineered proteins (i.e., self-assembly peptide-based hydrogels). For example, a commercially available hydrogel includes BD™ PuraMatrix™. BD™ PuraMatrix™ Peptide Hydrogel is a synthetic matrix that is used to create defined three dimensional (3D) micro-environments for cell culture.

For example, the hydrogel is a biocompatible polymer matrix that is biodegradable in whole or in part. Examples of materials which can form hydrogels include alginates and alginate derivatives, polylactic acid, polyglycolic acid, poly(lactic-co-glycolic acid) (PLGA) polymers, gelatin, collagen, agarose, natural and synthetic polysaccharides, polyamino acids such as polypeptides particularly poly(lysine), polyesters such as polyhydroxybutyrate and poly-epsilon.-caprolactone, polyanhydrides; polyphosphazines, poly(vinyl alcohols), poly(alkylene oxides) particularly poly(ethylene oxides), poly(allylamines)(PAM), poly(acrylates), modified styrene polymers such as poly(4-aminomethylstyrene), pluronic polyols, polyoxamers, poly(uronic acids), poly(vinylpyrrolidone), and copolymers of the above, including graft copolymers. Synthetic polymers and naturally-occurring polymers such as, but not limited to, collagen, fibrin, hyaluronic acid, agarose, and laminin-rich gels may also be used.

A preferred material for the hydrogel is alginate or modified alginate material. Alginate molecules are comprised of (1-4)-linked β-D-mannuronic acid (M units) and a L-guluronic acid (G units) monomers, which can vary in proportion and sequential distribution along the polymer chain. Alginate polysaccharides are polyelectrolyte systems which have a strong affinity for divalent cations (e.g., Ca⁺², Mg⁺², Ba⁺²) and form stable hydrogels when exposed to these molecules.

The compositions described herein are suitable for clinical use, e.g., bone repair, regeneration, or formation; muscle repair, regeneration, or formation; and dermal repair, regeneration, or formation. For example, the compositions are applied to bone fractures alone or together with a bone adhesive (cement) or glue or to diseased or injured muscle tissue. The hydrogels (seeded with cells or without cells) are injected at the site of disease, injury, or fracture (in the case of bone or cartilage). For example, the hydrogels are injected into or onto bone. Exemplary bioactive factors for use in promoting bone or cartilage repair, regeneration, or formation include BMP-2, BMP-4, or RunX.

In some cases, the composition recruits cells to promote bone or cartilage repair, regeneration, or formation. Alternatively, the first hydrogel or the second hydrogel comprises an isolated bone cell selected from the group consisting of an osteoblast, an osteocyte, an osteoclast, and an osteoprogenitor. Alternatively, the first hydrogel or the second hydrogel comprises an isolated cartilage cell, wherein the isolated cartilage cell comprises a chondroblast. The isolated bone cell or an isolated cartilage cell is an autologous cell or an allogeneic cell.

For muscle applications, e.g., muscle tears, muscle strains, or muscle pulls, the hydrogels (seeded with or without cells) are injected at the site of injury. Suitable compositions for muscle applications include a composition comprising a first hydrogel and a second hydrogel, wherein the first hydrogel degrades at least 10% faster than the second hydrogel, and wherein the first hydrogel or the second hydrogel comprises a bioactive factor for use in muscle repair, regeneration, or formation. For example, the bioactive factor comprises MyoD.

In some cases, the composition recruits cells to promote muscle or cartilage repair, regeneration, or formation. Alternatively, the first hydrogel or the second hydrogel comprises an isolated muscle cell selected from the group consisting of a skeletal muscle cell, a cardiac muscle cell, a smooth muscle cell, and a myo-progenitor cell. The isolated muscle cell is an autologous cell or an allogeneic cell.

For dermal applications, e.g., burns, abrasions, lacerations, or disease, the hydrogels (seeded with cells or without cells) are applied directly to the site as poultice or wound dressing. Preferably, the majority of porogens (e.g., more than 50%, more than 60%, more than 70%, more than 80%, or more than 90%) within the bulk hydrogels are directed toward the skin surface and into the skin tissue when applied directly to the site (e.g., the burn). In this manner, the bioactive factors or cells are released into the surface of the skin or lower layers of the skin, and do not migrate away from the skin or target tissue. An exemplary bioactive factor for use in skin repair, regeneration, or formation is FGF.

In some cases, the composition recruits cells to promote skin or cartilage repair, regeneration, or formation. Alternatively, the first hydrogel or the second hydrogel comprises an isolated skin cell selected from the group consisting of a fibroblast, a dermal cell, an epidermal cell, or a dermal progenitor cell. The isolated skin cell is an autologous cell or an allogeneic cell.

Bioactive factors such as polynucleotides, polypeptides, or other agents are purified and/or isolated. Specifically, as used herein, an “isolated” or “purified” nucleic acid molecule, polynucleotide, polypeptide, or protein, is substantially free of other cellular material, or culture medium when produced by recombinant techniques, or chemical precursors or other chemicals when chemically synthesized. Purified compounds are at least 60% by weight (dry weight) the compound of interest. Preferably, the preparation is at least 75%, more preferably at least 90%, and most preferably at least 99%, by weight the compound of interest. For example, a purified compound is one that is at least 90%, 91%, 92%, 93%, 94%, 95%, 98%, 99%, or 100% (w/w) of the desired compound by weight. Purity is measured by any appropriate standard method, for example, by column chromatography, thin layer chromatography, or high-performance liquid chromatography (HPLC) analysis. A purified or isolated polynucleotide (ribonucleic acid (RNA) or deoxyribonucleic acid (DNA)) is free of the genes or sequences that flank it in its naturally-occurring state. Purified also defines a degree of sterility that is safe for administration to a human subject, e.g., lacking infectious or toxic agents.

Similarly, by “substantially pure” is meant a nucleotide or polypeptide that has been separated from the components that naturally accompany it. Typically, the nucleotides and polypeptides are substantially pure when they are at least 60%, 70%, 80%, 90%, 95%, or even 99%, by weight, free from the proteins and naturally-occurring organic molecules with they are naturally associated.

An “isolated nucleic acid” is a nucleic acid, the structure of which is not identical to that of any naturally occurring nucleic acid, or to that of any fragment of a naturally occurring genomic nucleic acid spanning more than three separate genes. The term covers, for example: (a) a DNA which is part of a naturally occurring genomic DNA molecule, but is not flanked by both of the nucleic acid sequences that flank that part of the molecule in the genome of the organism in which it naturally occurs; (b) a nucleic acid incorporated into a vector or into the genomic DNA of a prokaryote or eukaryote in a manner, such that the resulting molecule is not identical to any naturally occurring vector or genomic DNA; (c) a separate molecule such as a cDNA, a genomic fragment, a fragment produced by polymerase chain reaction (PCR), or a restriction fragment; and (d) a recombinant nucleotide sequence that is part of a hybridgene, i.e., a gene encoding a fusion protein. Isolated nucleic acid molecules according to the present invention further include molecules produced synthetically, as well as any nucleic acids that have been altered chemically and/or that have modified backbones.

A small molecule is a compound that is less than 2000 daltons in mass. The molecular mass of the small molecule is preferably less than 1000 daltons, more preferably less than 600 daltons, e.g., the compound is less than 500 daltons, 400 daltons, 300 daltons, 200 daltons, or 100 daltons.

The transitional term “comprising,” which is synonymous with “including,” “containing,” or “characterized by,” is inclusive or open-ended and does not exclude additional, unrecited elements or method steps. By contrast, the transitional phrase “consisting of” excludes any element, step, or ingredient not specified in the claim. The transitional phrase “consisting essentially of” limits the scope of a claim to the specified materials or steps “and those that do not materially affect the basic and novel characteristic(s)” of the claimed invention.

Other features and advantages of the invention will be apparent from the following description of the preferred embodiments thereof, and from the claims. Unless otherwise defined, all technical and scientific terms used herein have the same meaning as commonly understood by one of ordinary skill in the art to which this invention belongs. Although methods and materials similar or equivalent to those described herein can be used in the practice or testing of the present invention, suitable methods and materials are described below. All published foreign patents and patent applications cited herein are incorporated herein by reference. Genbank and NCBI submissions indicated by accession number cited herein are incorporated herein by reference. All other published references, documents, manuscripts and scientific literature cited herein are incorporated herein by reference. In the case of conflict, the present specification, including definitions, will control. In addition, the materials, methods, and examples are illustrative only and not intended to be limiting.

BRIEF DESCRIPTION OF THE DRAWINGS

FIGS. 1A-1I is a series of schematics, photographs, bar graphs, and line graphs showing the formation of pores in situ within hydrogels as demonstrated via imaging and mechanical properties testing. FIG. 1A is a schematic illustrating the formation of hydrogels. Left: micro-beads comprised of rapidly degradable hydrogels (red spheres) and mesenchymal stem cells (MSCs; green). Middle: micro-beads and MSCs are mixed with a second hydrogel forming polymer material (“bulk gel;” gray), which is crosslinked around the beads. Right: after degradation of the micro-beads in situ, an intact hydrogel network remains with a network of pores through which MSCs are released. FIG. 1B is a fluorescence micrograph of fluorescein (green) labeled porogens immediately after fabrication. FIG. 1C is a fluorescence micrograph of fluorescein (green) labeled porogens after encapsulation into an alginate hydrogel network. FIG. 1D is a bar graph illustrating elastic modulus measurements of either standard hydrogels (blue bars) or pore-forming hydrogels (red bars). At day 0, there is no statistically significant difference in the overall rigidity of pore-forming composites and standard hydrogels because pores have not formed; however, after 4 days, the modulus of the composite drops substantially because of pore formation. FIG. 1E is a series of scanning electron micrographs depicting pore-forming hydrogels immediately after formation (Day 0) showing a grossly intact network, 5 days after fabrication, or 10 days after fabrication, at which time significant pore formation was observed. FIG. 1F and FIG. 1G verify that the elastic modulus of the composite material (50% porogen volume fraction) is not substantially different from the elastic modulus of a standard hydrogel (no porogens), but that as voids form, the modulus of the composite drops substantially. The decrease in composite elastic modulus at one week corresponds to the density of voids, and at low porogen density, there is a linear relationship between the density of voids and decrease in composite elastic modulus. FIG. 1H and FIG. 1I illustrate that the fracture toughness of the composite material (25% porogen volume fraction) is initially similar to the fracture toughness of a standard hydrogel with no porogen, but shortly decreases to a fraction of the initial value. As with elastic modulus, the decrease in fracture toughness scales with the density of porogen, though in a non-linear manner. Scale bars: (B, C, and E): 1 mm.

FIGS. 2A-2H is a series of photographs, line graphs, and schematics illustrating mesenchymal stem cell deployment in-vitro. Specifically, FIGS. 2A-2H demonstrates stem cell release from pore-forming hydrogels in vitro, which also shows that release can be tuned by varying the composition of porogens and the compartmentalization of cells within porogens versus bulk. FIG. 2A is a line graph showing the kinetics of interconnected void formation assessed with a capillary assay (error bars are standard error of the mean, n=3-4). Interconnected voids formed over the first 7 days unless a very high fraction of porogens (above the percolation limit; 80%) were present. No substantial interconnected void formation was observed with a sub-percolation porogen density. FIG. 2B is a series of schematics and photographs showing 3-dimensional reconstructions of Calcein-AM stained cells distributed throughout pore-forming hydrogels. The substantial changes in cell morphology depict the cells' ability to migrate and proliferate within pore-forming hydrogels, whereas cells remained sparse and rounded within standard hydrogels. Ki-67 immunofluorescence (green) indicates increased proliferation, while nuclear counterstain (Hoescht, blue) higher cellularity, within pore-forming hydrogels compared to standard hydrogels. FIG. 2C is a set of fluorescence micrographs showing the effects of void formation on cellularity and cell morphology within pore forming hydrogels. Specifically, fluorescence micrographs of pore-forming hydrogels were stained for live mesenchymal stem cells (MSC) (Calcein-AM, green) or dead cells (Ethidium Homodimer, red) after 4-10 days in vitro. The spherical cell morphology denotes cells confined in a nanoporous material, and is present at short time frames in both materials, but only in standard hydrogels over longer time frames. FIG. 2D is a line graph showing the cumulative number of MSC released after 12 days as a function of porogen volume density. FIG. 2E is a line graph and schematic illustrating the kinetics of deployment for MSC encapsulated either into the bulk phase, the porogen phase of pore-forming scaffolds, or into standard hydrogels. Porogens were prepared with 7.5% oxidized alginate and crosslinked in 100 mM CaCl₂. FIG. 2F is a line graph showing the kinetics of MSC deployment from the porogen phase of pore-forming hydrogels as a function of porogen fabrication conditions for D1 cells encapsulated into porogens. FIG. 2G is a bar graph showing that quantitative analysis of ³H-thymidine incorporation indicates enhanced cell proliferation, in an RGD-dependent manner. FIG. 2H is a series of line graphs showing the effects of the degree of alginate oxidation degree on cell release. At a constant level of calcium to crosslink porogens (100 mM), increasing the degree of oxidation from 3-7.5% substantially increased the overall number of release cells, whereas lowered the degree of oxidation slightly delayed cell release. At a constant degree of porogen degree of oxidation (7.5%), increasing the concentration of calcium used to crosslink porogens from 25-100 mM lowered the overall number of released cells and slightly delayed the onset of cell release. Scale bars: (A): 100 μm.

FIGS. 3A-3D is a series of photographs and line graphs showing the results of controlling mesenchymal stem cell deployment, engraftment and proliferation in vivo. Specifically, FIGS. 3A-3D demonstrates stem cell release from pore-forming hydrogels in vivo within the subcutaneous space of nude mice. FIG. 3A is a photograph showing representative images of Nude mice into which 2×10⁶ mCherry-expressing MSC were deployed, either 7 or 30 days after injection in standard hydrogels (left), pore-forming hydrogels in which porogens were crosslinked with either 100 mM CaCl₂ (center), or saline (right). The bulk component of hydrogels was modified with 2 RGD/polymer chain. Initially, more cells engrafted in the saline only condition, but at later time-points, there were fewer cells in this condition and substantially more cells eventually engrafting when released from pore-forming hydrogels. FIG. 3B is a line graph showing the quantification of the relative radiant efficiency (proportional to cell density) of mCherry-MSC injected in pore-forming hydrogels, standard hydrogels, or saline. FIG. 3C is a line graph showing that decreasing the density of calcium used to crosslink porogens substantially decreased the overall density of released cells, and slightly delayed the kinetics of deployment (error bars are SEM, n=4-6). FIG. 3D is a series of photographs showing the ability of pore-forming hydrogels to enhance human mesenchymal stem cell mediated bone regeneration using a Nude Rat cranial defect model. Critically-sized defects were formed in the crania of Nude Rats (Charles River). Immediately after defect formation, commercially available human mesenchymal stem cells (Lonza) were transplanted into the defect space, either within saline (“Cells Only”), a standard hydrogel (2 RGD/alginate polymer, 60 kPa), or a Pore-Forming Hydrogel. Representative cross-sections of micro-computed tomographic analysis of new bone formation within cranial defects 4 weeks after implantation. Substantially more new bone forms within defects in which cells were delivered via pore-forming hydrogels.

FIGS. 4A-4D is a series of photomicrographs depicting the use of pore-forming hydrogels to release distinct populations at different times. FIG. 4A-FIG. 4C are fluorescent micrographs of GFP-expressing myoblasts and outgrowth endothelial cells (OECs) adherent to tissue culture plastic after 4 days of culture within pore-forming hydrogels in which the chemistry used to form porogens was varied and the different cell types were initially placed into distinct compartments: (FIG. 4A): myoblasts in bulk component, OECs in porogen component; (FIG. 4B): myoblasts in bead component, OECs in porogen component; (FIG. 4C): both myoblasts and OECs in bulk component. FIG. 4D is a representative micrograph of a plastic substrate on which equal numbers of GFP-myoblasts and OECs were seeded. Myoblasts outgrew OECs. Cells were stained with Ethidium Homodimer (red), so that GFP-myoblasts appear yellow and OECs appear red. Images taken at 10× magnification.

FIGS. 5A-5D is a series of photomicrographs depicting using pore-forming hydrogels for chemokine-mediated cell recruitment. Specifically, FIGS. 5A-5D shows the controlling chemokine-mediated cell recruitment by pore-forming hydrogels in vivo. Alginate was first oxidized and then reduced with sodium borohydride to make alcohol groups that replace what were originally sugars. FIG. 5A and FIG. 5B are fluorescent micrographs of dendritic cell recruitment into (FIG. 5A) standard, injectable alginate gel, and (FIG. 5B) pore-forming hydrogel. Both sets of hydrogels were loaded with 2 pg of granulocyte-macrophage colony stimulating factor. FIG. 5C and FIG. 5D are fluorescent micrographs of dendritic cell recruitment into pore-forming hydrogels fabricated with (FIG. 5C) oxidized or (FIG. 5D) reduced porogens. No chemokine was added. For histology, dendritic cells are stained for CD11c (green) and NIHC-II (red), with Hoescht nuclear counterstain (blue). In FIG. 5C and FIG. 5D, only nuclear staining (white) was performed. This figure shows the difference in host cell recruitment by materials with porogens formed from the oxidized alginate vs. reduced alginate. Scale bars: 100 pm.

FIGS. 6A-6F is a series of line graphs, bar charts, and photographs demonstrating the control of stem cell proliferation within pore-forming hydrogels, deployment from hydrogels, and ability to regenerate bone by varying bulk phase composition. FIG. 6A and FIG. 6B are line graphs showing the analysis of 24 hr 3H-thymidine incorporation (proportional to DNA synthesis) by mesenchymal stem cells (D1; red curve) or cumulative MSC deployment (blue curves) from pore-forming hydrogels after 7 days of culture as a function of (FIG. 6A) density of RGD peptides in bulk gels with 60 kPa modulus, or (FIG. 6B) elastic modulus of bulk hydrogels presenting 10 RGD peptides/alginate polymer (data are mean+/−SEM, n=3-5). RGD density had significant effects of cell proliferation, whereas elastic modulus had effects on both proliferation and release (p<0.05, ANOVA). FIG. 6C is a bar graph showing the analysis of DNA synthesis as a function of pore formation. FIG. 6D is a series of photographs showing staining for Ki-67 (proliferation marker, green) in D1 cells in cryosectioned pore-forming hydrogels after 50 days of culture. FIGS. 6E and 6F is a line graph showing the control of mesenchymal stem cell deployment, engraftment and proliferation in vivo. FIG. 6E is a line graph showing the quantification of the relative radiant efficiency (proportional to cell density) of mCherry-MSC injected in pore-forming hydrogels in which the bulk phase was modified with either 2 (♦), or 10 (▪) RGD peptides per alginate polymer. Alternatively, cells were injected in a standard hydrogel with 2 RGD peptides/alginate polymer (▴). FIG. 6F shows the quantification of the percentage of healing (new bone formation due to human MSC transplanted into Nude Rat cranial defects) as a function of the method of MSC delivery. Error bars are SEM, n=4-6.

FIGS. 7A-7C is a series of line graphs and bar frequency charts showing the mechanical properties and in-vitro degradation of hydrogels formed from binary alginates. FIG. 7A and FIG. 7B show elastic moduli (FIG. 7A) and degradation (FIG. 7B) of bulk hydrogels formed by crosslinking binary combinations of oxidized alginate (5% theoretical degree of oxidation) at a constant density of 20 mg/mL with unmodified, high M_(w) alginate. Degradation was assessed by comparing the dry mass after 4 days in-vitro to initial dry mass. FIG. 7C is a histogram of diameters of porogens formed from binary mixtures of 20 mg/mL oxidized alginate with 7.5 mg/mL unmodified alginate. Porogen diameter was measured by processing fluorescent micrographs of porogens prepared from aminofluorescenin-labeled alginates. Error bars are SD, n=3-4.

FIG. 8 is a schematic of an implantable biomaterial that mimics certain aspects of stem-cell niches in that it activates transplanted progenitor cells to proliferate and programs them to differentiate into cells that migrate into damaged tissues to participate in regeneration.

FIG. 9 is a schematic of a hydrogel scaffold (left, top), which controls transplanted cell fate through presentation of specific cues, but prevents transplanted cells from migrating out of, and host cells from migrating into, the material. Bottom: example data depicting the ability of a nanoporous hydrogel to control mesenchymal stem cell fate, in this case via elastic modulus. Right, top: schematic of a macroporous sponge which controls transplanted fate through presentation of specific cues, while also allowing host cells to migrate into, and transplanted cells to migrate out of the material.

FIG. 10 is a series of images depicting an alternative strategy to produce macroporous hydrogels. As described in the schematics (center, right), porogens are embedded into a “bulk” hydrogel, or photolithographic techniques are applied, to provide non-crosslinked regions of the bulk hydrogel. After crosslinking the bulk hydrogel, the non-crosslinked portions of the hydrogel and porogens are removed using solvents such as acetone. Left: image of a macroporous hydrogel.

FIG. 11 is a schematic and a bar chart depicting a strategy to create rapidly degrading alginate-based hydrogel porogens. Top left: chemical reaction scheme to oxidize alginate to alginate dialdehyde with NaIO₄. Top right: schematic depicting the loss of crosslinkable, guluronic-acid rich portions of alginate (short, straight segments), and the overall decrease in polymer M_(w) due to sodium periodate oxidation. Bottom: data depicting loss in dry mass over time from hydrogels made from 20 mg/mL non-modified alginate (squares), 20 mg/mL alginate dialdehyde (5% degree of oxidation; diamonds) or a binary mixture of 20 mg/mL alginate dialdehyde with 7.5 mg/mL unmodified alginate (triangles).

FIG. 12 is a schematic illustrating porogen fabrication and characterization.

FIG. 13 is a schematic showing the control of host cell recruitment with pore forming hydrogels. Specifically, shown in this Figure is a schematic of an implantable biomaterial system that mimics the icroenvironment of an infection, allowing the recruitment, programming and subsequent targeting of activated antigen-presenting dendritic cells to the lymph nodes to participate in a potent antitumour response.

FIG. 14 is a series of photomicrographs showing how the polymers used to comprise the porogen phase affect the degradation and processing of porogens. Specifically, shown in this Figure are fluorescence micrographs of porogens formed using fluorescein-labeled alginate dialdehyde. Immediately after crosslinking in 100 mM CaCl₂ (top), porogens are grossly intact, whether made using 20 mg/mL alginate dialdehyde (top left) or a binary mixture of 20 mg/mL alginate dialdehyde with 7.5 mg/mL unmodified alginate. However, after processing steps used to purify porogens and remove excess CaCl₂, porogens made with purely alginate dialdehyde were damaged, resulting in substantial change in morphology and release of fluorescein-labeled polymers into solution to yield a substantial level of background fluorescein fluorescence (bottom left). In contrast, binary mixtures of 20 mg/mL alginate dialdehyde with 7.5 mg/mL unmodified alginate resulted in porogens that could withstand processing steps.

DETAILED DESCRIPTION OF THE INVENTION

Over the recent decades, biocompatible polymers have been used to form scaffolds that act as carriers for cell transplantation, or to recruit host cell populations into the device. Generally, sponges such as poly(lactide-co-glycolide) (PLGA), or synthetic hydrogels such as alginate are used. However, both sets of materials have disadvantages. For example, sponges typically adsorb serum proteins, so it is difficult to control presentation of adhesive proteins or peptides (for example, RGD) from the material. Sponge materials also typically are not amenable to injection, and require an invasive surgery for implantation, and also expose transplanted or host cells to a host environment that may initially be hostile (for example, neutrophils present during inflammation may attack stem cells). On the other hand, synthetic hydrogels are typically injectable, allowing minimally-invasive delivery, and do not interact with proteins. However, prior to the invention described herein, the pore size in hydrogels was typically much smaller than the diameter of a eukaryote cell, making it difficult to expand a transplanted cell population, release transplanted cells to allow them to repair damaged tissues, or recruit host cells into the device.

The present invention comprises a method to form pores in situ within hydrogels following hydrogel injection. Pores form in situ via degradation of sacrificial porogens encapsulated within the surrounding hydrogel. The kinetics and onset of pore formation are controlled by manipulating material used to form porogens, and cells are encapsulated either into the porogens themselves or the hydrogel surrounding them. Examples demonstrate in vitro deployment, proliferation, and differentiation of stem cells, as well as in vivo stem cell deployment and chemokine-mediated cell recruitment. The system mediates controlled deployment of cells out of, or local recruitment of cells into, a polymer matrix via formation of pores within this matrix. The size, distribution, and formation kinetics of the pores are pre-determined by the user, while the integrity of the matrix surrounding pores, along with cells or biological factors inside this matrix, are unchanged.

Accordingly, described herein is the use of insoluble cues such as hydrogel adhesion ligand presentation and/or elastic modulus (i.e., stiffness) to generate materials which are 1) injectable; 2) allow the user to control cell fate using insoluble cues; and 3) form pores over time to deploy or recruit cells. Specifically, the methods described herein create pore-forming hydrogels, using a process that allows cells to be encapsulated into either the pore-forming phase (hereafter referred to as “porogen”) or the non- or slowly-degrading phase (hereafter referred to as “bulk”).

The invention provides methods for a generalized approach to create pore-forming hydrogels that allow cell encapsulation, and a means to control the kinetics of cell deployment out of, or recruitment into, the hydrogel. Hydrogel micro-bead “porogens” are formed, and are next encapsulated into a second, “bulk” hydrogel. The composition of polymers used to form porogen and bulk hydrogels may be varied; however, the porogen must degrade more rapidly (e.g., 10%, 20%, 50%, 2×, 5×, 10×, 20× or faster) than the bulk hydrogel. Cells or bioactive factors (e.g., growth factors such as granulocyte/macrophage colony stimulating factor (GM-CSF), vascular endothelial growth factor (VEGF), condensed oligonucleotides, e.g., CpG, or plasmid DNA) are optionally encapsulated either into the porogen phase, bulk hydrogel phase, or into both phases. The porogens degrade in situ over a time-course pre-determined by the user, at which point cells are released, or may migrate into the material. However, because they initially lack pores, pore-forming hydrogels are useful to provide mechanical support immediately after formation (FIGS. 1A-1I).

Cellular release or recruitment is manipulated by controlling the kinetics of porogen degradation. For example, the alginate polymers are oxidized to produce alginate dialdehyde, and the total number of cells released increases as the extent of oxidation increases (FIGS. 2A-2H, FIGS. 3A-3D). Alternatively, conditions used to crosslink the porogens are altered to manipulate the time at which significant porogen degradation and cell release begin to occur (FIGS. 2A-2H). Porogen chemistry can further be varied to facilitate, or inhibit, interaction with host proteins (FIGS. 5A-5D).

Cell release and cell fate are controlled by manipulating the biophysical and biochemical properties (e.g. elastic modulus and density of integrin-binding adhesion peptides such as RGD) of the bulk hydrogel. For example, pore formation, bulk hydrogel RGD density and bulk hydrogel elasticity all affect cell proliferation within these materials (FIGS. 2A-2H, FIGS. 6A-6F). The orthogonal processing of porogens and bulk separate from one another enhances the ability to tune the system to manipulate cell release and cell fate. For example, stem cell lineage commitment is modulated by varying elastic modulus or RGD density, independent of the kinetics of pore formation. In contrast, other techniques used to form macro-porous materials (e.g. solvent-based extraction of porogens) are not compatible with cell encapsulation, and typically affect the physical properties of both the porogen and bulk phases. The physical and biochemical properties rapidly degrading bulk hydrogel materials change continuously over the course of degradation. These parameters are harnessed to design the bulk phase to regulate cell fate.

Hydrogel Compositions

Hydrogels comprise a network of polymer chains that are hydrophilic. Hydrogel (also called aquagel) is sometimes found as a colloidal gel in which water is the dispersion medium. Hydrogels are highly absorbent (they can contain over 99.9% water) natural or synthetic polymers. Hydrogels also possess a degree of flexibility very similar to natural tissue, due to their significant water content. Hydrogel is comprised of cross-linked polymers. Exemplary hydrogels are comprised materials that are compatible with cell encapsulation such as alginate, polyethylene glycol (PEG), PEG-acrylate, agarose, and synthetic protein (e.g., collagen or engineered proteins (i.e., self-assembly peptide-based hydrogels). For example, a commercially available hydrogel includes BD™ PuraMatrix™ Peptide Hydrogel, which is a synthetic matrix that is used to create defined three dimensional (3D) micro-environments for cell culture.

For example, the hydrogel is a biocompatible polymer matrix that is biodegradable in whole or in part. Examples of materials which can form hydrogels include alginates and alginate derivatives, polylactic acid, polyglycolic acid, poly(lactic-co-glycolic acid) (PLGA) polymers, gelatin, collagen, agarose, natural and synthetic polysaccharides, polyamino acids such as polypeptides particularly poly(lysine), polyesters such as polyhydroxybutyrate and poly-epsilon.-caprolactone, polyanhydrides; polyphosphazines, poly(vinyl alcohols), poly(alkylene oxides) particularly poly(ethylene oxides), poly(allylamines)(PAM), poly(acrylates), modified styrene polymers such as poly(4-aminomethylstyrene), pluronic polyols, polyoxamers, poly(uronic acids), poly(vinylpyrrolidone), and copolymers of the above, including graft copolymers. Synthetic polymers and naturally-occurring polymers such as, but not limited to, collagen, fibrin, hyaluronic acid, agarose, and laminin-rich gels may also be used.

A preferred material for the hydrogel is alginate or modified alginate material. Alginate molecules are comprised of (1-4)-linked β-D-mannuronic acid (M units) and a L-guluronic acid (G units) monomers, which can vary in proportion and sequential distribution along the polymer chain. Alginate polysaccharides are polyelectrolyte systems which have a strong affinity for divalent cations (e.g., Ca⁺², Mg⁺², Ba⁺²) and form stable hydrogels when exposed to these molecules.

Synthetic hydrogels are typically injectable, allow for minimally-invasive delivery, and do not interact with proteins. Hence, the presentation of adhesion proteins or peptides is precisely controlled. Moreover, synthetic hydrogels typically have a pore mesh size that is much smaller than cells (<10 nm, whereas cells are >10 um), which prevents host cells from attacking transplanted cells. However, this small pore size also prevents transplanted cells from proliferating extensively within the material, and also precludes their eventually being released to affect various functions (for example, regeneration of functional tissue or destruction of diseased tissue).

Several techniques have been introduced to combine desirable features of hydrogels and sponges—for example, rigid microspheres have been encapsulated into hydrogels, and then extracted with solvents (e.g. acetone) to leave behind a macroporous hydrogel, and freeze-drying has been applied to generate macroporous hydrogels. Hydrogels can be modified to rapidly degrade in vivo to release host cells. However, prior to the invention described herein, none of the approaches allowed for the combination of a non-degrading (or slowly degrading) material component with cell encapsulation. The mechanical properties and biochemical composition of hydrogel materials strongly affect cell fate, and degradation in-and-of itself may intrinsically regulate cell fate.

Pore-Forming Compositions

Hydrogel micro-beads (“porogens”) are formed. Next, porogens are encapsulated into a “bulk” hydrogel that is either non-degradable or which degrades at a slow rate compared to the porogens. Cells are optionally encapsulated either into the porogen or bulk compartment. Immediately after hydrogel formation, or injection into the desired site in vivo, the composite material lacks pores, and serves as a surgical bulking agent. Subsequently, porogen degradation causes pores to form in situ, and encapsulated cells deploy away from the composite material and into surrounding tissues or remote tissues, e.g., lymph nodes, in the body. The size and distribution of pores are controlled during porogen formation, and mixing with the polymers which form the bulk hydrogel.

Alternatively, the hydrogel is injected without encapsulated cells, and pore formation is used as a means of recruiting host cells, in combination or independent of chemokines released from either the bulk or porogen component. The porogens are comprised of any biocompatible polymer, as long as they degrade more rapidly than the material used to form the “bulk” hydrogel, and are initially mechanically stable enough to withstand being mixed with the polymer which forms the bulk hydrogel phase. The “bulk” is comprised of any hydrogel-forming polymer.

Alginate Compositions

The polymers utilized in the compositions and methods are naturally-occurring or synthetically made. In one example, both the porogens and bulk hydrogels are formed from alginate. “Alginate” as that term is used here, refers to any number of derivatives of alginic acid (e.g., calcium, sodium or potassium salts, or propylene glycol alginate). See, e.g., PCT/US97/16890, hereby incorporated by reference.

The alginate polymers suitable for porogen formation have a Dalton molecular weight from 5,000 to 500,000 Da. The polymers are optionally further modified (e.g., by oxidation with sodium periodate, (Bouhadir et al., 2001, Biotech. Prog. 17:945-950, hereby incorporated by reference), to facilitate rapid degradation. In the examples described below, the polymers were crosslinked by extrusion through a nebulizer with co-axial airflow into a bath of divalent cation (for example, Ca2+ or Ba2+) to form hydrogel micro-beads. The higher the airflow rate, the lower the porogen diameter.

The concentration of divalent ions used to form porogens may vary from 5 to 500 mM, and the concentration of polymer from 1% to 5% by weight. However, any method which produces porogens that are significantly smaller than the bulk phase is suitable. Porogen chemistry can further be manipulated to produce porogens that have a some interaction with host proteins and cells (e.g., alginates oxidized to an extent of >5% of sugar resides interact significantly with host cells, FIGS. 5A-5D), or to inhibit this interaction (e.g., oxidized alginates that are reduced with NaBH4 exhibit minimal interaction with protein or with host cells, FIG. 5A-5D).

The alginate polymers suitable for formation of the bulk hydrogel have a Dalton molecular weight from 5,000 to 500,000 Da. The polymers may be further modified (for example, by oxidation with sodium periodate), to facilitate degradation, as long as the bulk hydrogel degrades more slowly than the porogen. The polymers may also be modified to present biological cues to control cell responses (e.g., integrin binding adhesion peptides such as RGD). Either the porogens or the bulk hydrogel may also encapsulate bioactive factors such as oligonucleotides, growth factors or drugs to further control cell responses. The concentration of divalent ions used to form the bulk hydrogel may vary from 5 to 500 mM, and the concentration of polymer from 1% to 5% by weight. The elastic modulus of the bulk polymer is tailored, e.g., to control the fate of encapsulated cells.

Example 1: Forming Pores In Situ within Hydrogels

The formation of pores in situ within hydrogels as demonstrated via imaging and mechanical properties testing is shown in FIGS. 1A-1I. As shown in FIG. 1A, micro-beads comprised of rapidly degradable hydrogels (red spheres) were mixed with a second hydrogel forming polymer material, which is crosslinked around the beads. After degradation of the micro-beads in situ, an intact hydrogel network (pink) remained with a network of pores. The elastic modulus measurements of either standard hydrogels (left bars) or pore-forming hydrogels (right bars) were determined (FIG. 1D). At day 0, there was no statistically significant difference in the overall rigidity of pore-forming composites and standard hydrogels because pores have not formed; however, after 4 days, the modulus of the composite drops substantially because of pore formation.

Additional methods relevant to generating the hydrogels described herein are as follows. Bouhadir et al. Polymer 1999; 40: 3575-84 (incorporated herein by reference) describes the oxidation of alginate with sodium periodate, and characterizes the reaction. Bouhadir et al. Biotechnol. Prog. 2001; 17: 945-50 (incorporated herein by reference) describes oxidation of high molecular weight alginate to form alginate dialdehyde (alginate dialdehyde is high M_(w) alginate in which a certain percent, (e.g., 5%), of sugars in alginate are oxidized to form aldehydes), and application to make hydrogels degrade rapidly. Kong et al. Polymer 2002; 43: 6239-46 (incorporated herein by reference) describes the use of gamma-irradiation to reduce the weight-averaged molecular weight (M_(w)) of guluronic acid (GA) rich alginates without substantially reducing GA content (e.g., the gamma irradiation selectively attacks mannuronic acid, MA blocks of alginate). Alginate is comprised of GA blocks and MA blocks, and it is the GA blocks that give alginate its rigidity (elastic modulus). Kong et al. Polymer 2002; 43: 6239-46 (incorporated herein by reference) shows that binary combinations of high M_(w), GA rich alginate with irradiated, low M_(w), high GA alginate crosslinks with calcium to form rigid hydrogels, but which degrade more rapidly and also have lower solution viscosity than hydrogels made from the same overall weight concentration of only high M_(w), GA rich alginate. Alsberg et al. J Dent Res 2003; 82(11): 903-8 (incorporated herein by reference) describes degradation profiles of hydrogels made from irradiated, low M_(w), GA-rich alginate, with application in bone tissue engineering. Kong et al. Adv. Mater 2004; 16(21): 1917-21 (incorporated herein by reference) describes control of hydrogel degradation profile by combining gamma irradiation procedure with oxidation reaction, and application to cartilage engineering.

Techniques to control degradation of hydrogen biomaterials are well known in the art. For example, Lutolf M P et al. Nat Biotechnol. 2003; 21: 513-8 (incorporated herein by reference) describes poly(ethylene glycol) based materials engineered to degrade via mammalian enzymes (MMPs). Bryant S J et al. Biomaterials 2007; 28(19): 2978-86 (U.S. Pat. No. 7,192,693 B2; incorporated herein by reference) describes a method to produce hydrogels with macro-scale pores. A pore template (e.g., poly-methylmethacrylate beads) is encapsulated within a bulk hydrogel, and then acetone and methanol are used to extract the porogen while leaving the bulk hydrogel intact. Silva et al. Proc. Natl. Acad. Sci USA 2008; 105(38): 14347-52 (incorporated herein by reference; US 2008/0044900) describes deployment of endothelial progenitor cells from alginate sponges. The sponges are made by forming alginate hydrogels and then freeze-drying them (ice crystals form the pores). These materials improve the therapeutic effect of the cells (compared to cells delivered alone), but these materials must be implanted surgically (i.e., non-injectable), are not amenable to cell encapsulation (cells will die when freeze dried), and this strategy makes it difficult to control cell fate by controlling elastic modulus. Ali et al. Nat Mater 2009 (incorporated herein by reference) describes the use of porous scaffolds to recruit dendritic cells and program them to elicit anti-tumor responses. Huebsch et al. Nat Mater 2010; 9: 518-26 (incorporated herein by reference) describes the use of hydrogel elastic modulus to control the differentiation of encapsulated mesenchymal stem cells.

Described herein is the use of insoluble cues such as hydrogel adhesion ligand presentation and/or elastic modulus (i.e., stiffness) to generate materials which are 1) injectable; 2) allow the user to control cell fate using insoluble cues; and 3) form pores over time to deploy or recruit cells. Specifically, the methods described herein create pore-forming hydrogels, using a process that allows cells to be encapsulated into either the pore-forming phase (hereafter referred to as “porogen”) or the non- or slowly-degrading phase (hereafter referred to as “bulk”). In the methods described herein, the porogen degrades by hydrolysis rather than by solvents, which means that cells are encapsulated either into the porogen or the bulk gel around them, and there is very little chance that proteins or other bioactive compounds encapsulated into the gel would be denatured.

As described in detail below, porogens stayed intact during encapsulation, but rapidly degraded to yield voids that were visible by scanning electron microscopy, and resulted in loss of elastic modulus and fracture toughness of the composite materials. Specifically, scanning electron micrographs (SEM) showed that pore-forming hydrogels immediately after formation (Day 0) possessed a grossly intact network; however, by 10 days after fabrication, significant pore formation was observed (FIG. 1E). The elastic modulus of the composite material (50% porogen volume fraction) was not substantially different from the elastic modulus of a standard hydrogel (no porogens); however, as voids form, the modulus of the composite drops substantially (FIG. 1F and FIG. 1G). The decrease in composite elastic modulus at one week corresponds to the density of voids, and at low porogen density, there is a linear relationship between the density of voids and decrease in composite elastic modulus. The fracture toughness of the composite material (25% porogen volume fraction) was initially similar to the fracture toughness of a standard hydrogel with no porogen, but it decreases to a fraction of the initial value (FIG. 1H and FIG. 1I). As with elastic modulus, the decrease in fracture toughness scales with the density of porogen, though in a non-linear manner. These results demonstrate that porogens stay intact during encapsulation and degrade in situ to form voids.

Example 2: In Vitro and In Vivo Release of Cells

Pore-forming hydrogels were formed by encapsulating degradable alginate porogens, along with bone marrow stromal stem cells (D1) into high molecular weight bulk alginate gel. Porogens were formed with a binary mixture of 20 mg/mL of alginate dialdehyde (theoretical oxidation of 7.5% of alginate sugar residues in high Mw, high guluronic acid content alginate) and 7.5 mg/mL high Mw, high guluronic acid (GA) content alginate. This polymer mixture was extruded through a glass nebulizer with co-axial nitrogen airflow into a bath of 0.1M CaCl₂ and 0.1M HEPES to crosslink polymers. Porogens were washed extensively with serum free cell culture medium. The bulk hydrogel was formed by 20 mg/mL high Mw, high GA-content alginate modified with 2 RGD peptides per alginate polymer. D1 cells and porogens were mixed into the bulk hydrogel material using syringes and then the composite was crosslinked with Calcium Sulfate. The number of D1 cells released from this system over time in vitro is shown in FIG. 2A-2H. The kinetics of release could be modified by 1) controlling the concentration of CaCl₂ used to form porogens, by 2) varying the composition (degree of oxidation) of porogens, and by 3) varying the compartmentalization of cells (either within porogens or within bulk gel).

Specifically, mesenchymal stem cell deployment in-vitro is illustrated in FIGS. 2A-2H. Stem cells were released from pore-forming hydrogels in vitro, and this release can be tuned by varying the composition of porogens and the compartmentalization of cells within porogens versus bulk. The effects of porogen density (0-80 volume percent) on cellularity and efflux from pore-forming hydrogels are shown in FIG. 2C. Specifically, fluorescence micrographs of pore-forming hydrogels were stained for live mesenchymal stem cells (MSC) (Calcein-AM, green) or dead cells (Ethidium Homodimer, red) after 4-10 days in vitro. The spherical cell morphology denotes cells confined in a nanoporous material, and is present at short time frames in both materials, but only in standard hydrogels over longer time frames. The cumulative number of MSC released after 12 days as a function of porogen volume density is shown in FIG. 2D.

As shown in FIG. 2D, the size of the porogen is related to the size of the overall composite material. Specifically, for the material to stay intact, the porogen diameter is <10% of the smallest dimension of the overall composite. The density of porogens is between 10-80 percent of the overall volume for both cell recruitment and cell release, e.g., between 15% and 75%, between 20% and 70%, between 25% and 65%, between 30% and 60%, or between 35% and 55% of the overall volume. Preferably, the density of porogens is at least 50% of the overall volume.

Physical and in vitro studies were performed to determine the kinetics of interconnected void formation, and the corresponding kinetics of mesenchymal stem cell release (FIG. 2A and FIG. 2E, respectively. A clonally derived, commercially available mouse mesenchymal stem cell line (D1) was used for these in vitro studies. A capillary assay was used to assess void formation. Briefly, the density of interconnected voids was measured by first weighing buffer-saturated composite pore-forming gels, and then re-weighing gels after wicking away water by gently touching the surface of the gel with a paper towel. The void fraction was calculated based on the relative change in mass. For in vitro cell release assays, the bulk component of pore-forming hydrogels was modified with 2 RGD peptides/alginate polymer, and had an elastic modulus of 60 kPa. Interconnected voids formed over the first 7 days unless a very high fraction of porogens (above the percolation limit; 80%) were present. No substantial interconnected void formation was observed with a sub-percolation porogen density. The kinetics of MSC deployment from the porogen phase of pore-forming hydrogels as a function of porogen fabrication conditions for D1 cells encapsulated into porogens is illustrated in FIG. 2F.

Release studies were subsequently performed with a mouse MSC line. Cell release was observed in proportion to overall pore density and gradual change in cell morphology, reflecting a loss of micron-scale confinement. Experiments were performed to determine the effects of pore formation on cellularity and cell proliferation within hydrogels. Cellularity was determined qualitatively using Calcein-AM staining, while proliferation was determined qualitatively by immunostaining for Ki-67 expression or quantitatively by measuring ³H-thymidine incorporation. Three-dimensional reconstructions of Calcein-AM stained cells distributed throughout pore-forming hydrogels are presented in FIG. 2B. The substantial changes in cell morphology depict the cells' ability to migrate and proliferate within pore-forming hydrogels, whereas cells remained sparse and rounded within standard hydrogels. Ki-67 immunofluorescence indicated higher cellularity, and increased proliferation, within pore-forming hydrogels compared to standard hydrogels. Quantitative analysis of ³H-thymidine incorporation indicated enhanced cell proliferation in an RGD-dependent manner (FIG. 2G).

Studies were performed to determine whether varying the chemical composition or cross-linking conditions used to form porogens would modulate the kinetics of cell release (Bouhadir K H, Lee K Y, Alsberg E, Damm K L, Anderson K W, Mooney D J. Degradation of Partially Oxidized Alginate and Its Potential Application for Tissue Engineering. Biotechnol. Prog. 2001; 17: 945-50). FIG. 2H shows that cell release kinetics were controlled by the compartmentalization of cells either into porogens or the bulk gel surrounding them, and that porogens that were crosslinked with a lower concentration of calcium degraded more slowly, releasing cells at a later time point. The lower concentration of calcium used to fabricate porogens led to more homogeneous cross-linking.

Pore-forming hydrogels were formed with a constant bulk component (2 RGD/polymer, 60 kPa), and constant porogen density (50%), but varying porogen composition. The chemical composition of porogens was manipulated by varying the theoretical degree of oxidation of the alginate polymers. Oxidation degree was controlled by varying the ratio of sodium periodate to alginate during the oxidation reaction (Bouhadir 2001). Binary mixtures of 20 mg/mL oxidized alginate with 5 mg/mL unmodified, high M_(w) alginate, were used to form porogens. Porogens were formed by crosslinking in a bath of 25-100 mM CaCl₂. The effects of the degree of alginate oxidation degree on cell release are shown in FIG. 2H. At a constant level of calcium to crosslink porogens (100 mM), increasing the degree of oxidation from 3-7.5% substantially increased the overall number of release cells, whereas lowering the degree of oxidation slightly delayed cell release. At a constant degree of porogen degree of oxidation (7.5%), increasing the concentration of calcium used to crosslink porogens from 25-100 mM lowered the overall number of released cells and slightly delayed the onset of cell release.

Example 3: Controlling Mesenchymal Stem Cell Deployment, Engraftment, and Proliferation In Vivo

Finally, in vivo studies were performed to determine if pore-forming hydrogels could be used to manipulate the release kinetics of MSC in vivo. For this, mouse MSC expressing mCherry were transplanted subcutaneously into Nude mice. Cell engraftment, proliferation and deployment were observed with non-invasive fluorescence imaging. This showed that not only did pore-forming gels delay engraftment compared to cells delivered in saline, but that these materials ultimately led to more proliferation. The hydrogels provide a micro-environment ammenable to proliferation after pores have formed. Finally, as these materials were useful to promote MSC release and expansion in vivo, human MSC were administered to regenerate cranial defects on nude rats. This led to improved regeneration of mineralized bone, even at an early time-point.

Specifically, for in vivo studies, D1 cells were modified to constitutively express a detectable marker, e.g., mCherry or green fluorescent protein (GFP), and were encapsulated either into standard bulk gels with no porogens, pore-forming hydrogels, or mixed with saline. Next, cells were injected into the backs of Nude mice through 18-gauge needles. Cell release and proliferation over time were monitored via mCherry fluorescence observed on an IVIS system (Caliper Life Sciences). These data revealed significantly more cell release and proliferation from pore-forming hydrogels than from standard gels (FIGS. 3A-3D). Moreover, although cells did proliferate if delivered by simple saline injection, deployment from pore-forming hydrogels 1) altered the kinetics of local cell delivery and proliferation, and 2) eventually led to a substantially higher number of delivered cells (FIGS. 3A-3D).

Specifically, experiments were performed to determine the ability to manipulate the kinetics of cell release in vivo by varying the composition of porogens. As shown in FIGS. 3A-3D, stem cells were released from pore-forming hydrogels in vivo within the subcutaneous space of nude mice. 2×10⁶ mCherry-expressing MSC were deployed into Nude mice, either 7 or 30 days after injection in standard hydrogels (left), pore-forming hydrogels in which porogens were crosslinked with either 100 mM or 50 MM CaCl₂ (center), or saline (right). The bulk component of hydrogels was modified with 2 RGD/polymer chain. Initially, more cells engrafted in the saline only condition, but at later time-points, there were fewer cells in this condition and substantially more cells eventually engrafting when released from pore-forming hydrogels. The quantification of the relative radiant efficiency (proportional to cell density) of mCherry-MSC injected in pore-forming hydrogels, standard hydrogels, or saline is shown in FIG. 3B. Decreasing the density of calcium used to crosslink porogens substantially decreased the overall density of released cells, and slightly delayed the kinetics of deployment (FIG. 3C; error bars are SEM, n=4-6). The ability of pore-forming hydrogels to enhance human mesenchymal stem cell mediated bone regeneration was demonstrated in a Nude Rat cranial defect model (FIG. 3D). Critically-sized defects were formed in the crania of Nude Rats (Charles River). Immediately after defect formation, commercially available human mesenchymal stem cells (Lonza) were transplanted into the defect space, either within saline (“Cells Only”), a standard hydrogel (2 RGD/alginate polymer, 60 kPa), or a Pore-Forming Hydrogel. The top row of FIG. 3D shows representative cross-sections of micro-computed tomographic analysis of new bone formation within cranial defects 4 weeks after implantation. Substantially more new bone forms within defects in which cells were delivered via pore-forming hydrogels. Doxycycline incorporation (green) into newly-forming bone at 12 weeks after implantation demonstrates that pore-forming hydrogels lead to positive doxycycline staining within the tissue rather than false-positive staining of the subcutaneous tissues.

Example 4: In Vitro Release of Two Different Cell Populations at Distinct Times

Pore-forming hydrogels were formed as described in Examples 1 and 2. Equal numbers (approximately 106 cells/mL of composite pore-forming hydrogel) GFP-expressing myoblasts and outgrowth endothelial cells (OECs, vascular progenitor cells) were encapsulated into different compartments of the material. After 5 days of culture in vitro, cells that were released and adherent to tissue culture plastic were stained with Ethidium Homodimer (EtD-1; red). As shown in FIGS. 4A-4D, the cell type encapsulated into the bulk gel deployed more rapidly. This pattern of deployment occurred even for OECs, which migrate more slowly and proliferate less extensively than GFP-myoblasts (based on analysis of substrates onto which equal numbers of both cell types were added; FIG. 4D).

Specifically, the use of pore-forming hydrogels to release distinct populations at different times is shown in FIGS. 4A-4D. Fluorescent micrographs of GFP-expressing myoblasts and outgrowth endothelial cells (OECs) adherent to tissue culture plastic after 4 days of culture within pore-forming hydrogels in which the chemistry used to form porogens was varied and the different cell types were initially placed into distinct compartments are shown in FIGS. 4A-4C. FIG. 4A depicts myoblasts in bulk component, OECs in porogen component, while FIG. 4B depicts myoblasts in bead component, OECs in porogen component. FIG. 4C depicts both myoblasts and OECs in bulk component. Equal numbers of GFP-myoblasts and OECs were seeded onto a plastic substrate (FIG. 4D). Myoblasts outgrew OECs. Cells were stained with Ethidium Homodimer (red), so that GFP-myoblasts appear yellow and OECs appear red.

Example 5: Recruitment of Host Lymphocytes from Subcutaneous Tissues by Pore-Forming Hydrogels with Different Porogen Formulations

Pore-forming hydrogels were formed as described in Examples 1 and 2. To form the porogen phase, 7.5 mg/mL of high Mw, GA-rich alginate polymer was combined with 20 mg/mL of either alginate dialdehyde (7.5% theoretical degree of oxidation) or alginate dialdehyde in which aldehyde groups were reduced to alcohol groups. Pore-forming hydrogels without encapsulated cells were next injected into the backs of C57/BL6 or Balb/c mice. After 14 days, recruitment of host dendritic cells was observed by histology.

As described in detail below, pore-forming hydrogels were utilized for chemokine-mediated cell recruitment. Alginate was first oxidized and then reduced with sodium borohydride to make alcohol groups that replace what were originally sugars. FIG. 5A and FIG. 5B show a comparison of dendritic cell (DC) recruitment by a standard, degradable alginate hydrogel versus a pore-forming alginate hydrogel. Both sets of hydrogels were loaded with 2 ug of granulocyte-macrophage colony stimulating factor. This indicates substantially more infiltration of cells from the tissue adjacent the hydrogel (highly cellular area near the edge of the image) into the pore-forming hydrogel. FIG. 5C and FIG. 5D show a comparison of baseline DC invasion into pore-forming hydrogels in which porogens were formed from alginate dialdehyde (FIG. 5C) or from reduced alginate dialdehyde (FIG. 5D) without GM-CSF. Substantially less baseline cell infiltration occurred absent GM-CSF. For histology, dendritic cells are stained for CD11c (green) and NIHC-II (red), with Hoescht nuclear counterstain (blue). This figure shows the difference in host cell recruitment by materials with porogens formed from the oxidized alginate vs. reduced alginate.

Example 6: Control of Stem Cell Proliferation within Pore-Forming Hydrogels by Varying Bulk Phase Composition

The purpose of this approach is to manipulate cell expansion and release using insoluble cues. Thus, it was determined whether the density of adhesion ligands and mechanical properties of the non-degrading hydrogel surrounding porogens would have effects on the cells. As shown in FIGS. 6A-6F, the density of ligands significantly altered DNA synthesis, whereas altering elastic modulus altered both DNA synthesis and cell release over 1 week. Insoluble cues like adhesion ligand density had effects on cells over long time-frames, as shown by histology in FIG. 6D.

Specifically, studies were performed to determine whether the composition of the bulk component of pore-forming hydrogels could modulate cell proliferation and engraftment in vivo. The analysis of 24 hr ³H-thymidine incorporation (proportional to DNA synthesis) by mesenchymal stem cells (D1; red curve) or cumulative MSC deployment (blue curves) from pore-forming hydrogels after 7 days of culture as a function of density of RGD peptides in bulk gels with 60 kPa modulus, or elastic modulus of bulk hydrogels presenting 10 RGD peptides/alginate polymer (data are mean+/−SEM, n=3-5) is shown in FIG. 6A and FIG. 6B. RGD density had significant effects of cell proliferation, whereas elastic modulus had effects on both proliferation and release (p<0.05, ANOVA). The analysis of DNA synthesis as a function of pore formation is shown in FIG. 6C. Staining for Ki-67 (proliferation marker, green) in D1 cells in cryosectioned pore-forming hydrogels after 50 days of culture is shown in FIG. 6D.

Control of Deployed Stem Cell Fate Via Composition of the Bulk Hydrogel

Pore forming hydrogels were formed as described in Example 1. By manipulating the composition (density of integrin-binding RGD peptides and elastic modulus) of the bulk hydrogel, it was possible to control mesenchymal stem cell (MSC) proliferation and release in vitro. In vivo, the overall density of mCherry-labeled mouse MSC deployed into the subcutaneous space could be increased by increasing the density of RGD peptides from 2 to 10 peptides per alginate polymer chain (FIG. 6E). For therapeutic studies, human MSC were deployed into Nude Rat cranial defects. After 4 weeks, rats were euthanized, and the degree of healing (due to new bone formation) was assessed by Hematoxylin/Eosin staining. Briefly, the area of newly formed bone within the defect was divided by the total area of the defect to generate the “Percent Healing” metric. Using this quantitative metric, it was found that delivering MSC in pore-forming hydrogels was substantially better than delivery via standard hydrogels or saline in terms of ability to induce new bone formation (FIG. 6F). Moreover, the elastic modulus of the bulk hydrogel component had a substantial effect on new bone formation at 4 weeks, as deployment from a pore-forming hydrogel with a 60 kPa, 10 RGD/alginate polymer bulk phase led to significantly more bone formation (p<0.05, 2-tailed t-test) than deployment from a pore-forming hydrogel with an 8 kPa, 10 RGD/alginate polymer bulk phase.

Thus, when mCherry-labeled D1 were deployed into the subcutaneous tissues of Nude mice via pore-forming hydrogels, increasing the RGD density of the bulk component from 2 to 10 RGD peptides/alginate polymer substantially increased the overall number of engrafted cells without significantly affecting cell deployment kinetics.

Though the example here demonstrated an effect of bulk hydrogel elasticity on cell-mediated tissue regeneration, as described herein, many other aspects of the bulk hydrogel phase—for example, the presentation of matrix-bound growth factors or peptide-mimics thereof—are engineered to influence cell-mediated tissue regeneration.

Example 7: Mechanical Properties and In-Vitro Degradation of Hydrogels Formed from Binary Alginates

Elastic moduli and degradation of bulk hydrogels formed by cross-linking binary combinations of oxidized alginate (5% theoretical degree of oxidation) at a constant density of 20 mg/mL with unmodified, high M_(w) alginate are shown in FIG. 7A and FIG. 7B. Degradation was assessed by comparing the dry mass after 4 days in-vitro to initial dry mass. The diameters of porogens formed from binary mixtures of 20 mg/mL oxidized alginate with 7.5 mg/mL unmodified alginate is shown in FIG. 7C. Porogen diameter was measured by processing fluorescent micrographs of porogens prepared from aminofluorescenin-labeled alginates.

Other Embodiments

While the invention has been described in conjunction with the detailed description thereof, the foregoing description is intended to illustrate and not limit the scope of the invention, which is defined by the scope of the appended claims. Other aspects, advantages, and modifications are within the scope of the following claims.

The patent and scientific literature referred to herein establishes the knowledge that is available to those with skill in the art. All United States patents and published or unpublished United States patent applications cited herein are incorporated by reference. All published foreign patents and patent applications cited herein are hereby incorporated by reference. Genbank and NCBI submissions indicated by accession number cited herein are hereby incorporated by reference. All other published references, documents, manuscripts and scientific literature cited herein are hereby incorporated by reference.

While this invention has been particularly shown and described with references to preferred embodiments thereof, it will be understood by those skilled in the art that various changes in form and details may be made therein without departing from the scope of the invention encompassed by the appended claims. 

1.-25. (canceled)
 26. A composition comprising a composite scaffold composition, wherein the composite scaffold composition: (i) is a polymeric scaffold composition (ii) lacks macropores having a diameter of at least 20 μm; (iii) comprises a crosslinked bulk hydrogel encapsulating sacrificial porogen hydrogel micro-beads having a diameter between about 20 μm and about 500 and (iv) comprises sacrificial porogen hydrogel micro-beads at a density of between 50% to 80% of the overall volume of the composite polymeric composition, and that comprise oxidized alginate or a shorter polymer than said bulk hydrogel such that the sacrificial porogen hydrogel micro-beads degrade at least 10% faster than said bulk hydrogel in situ; wherein upon administration to a subject the sacrificial porogen hydrogel micro-beads degrade in situ to form a network of macropores having a diameter between about 20 μm and about 500 μm in their place, and an intact hydrogel network which allows for the recruitment of cells into the scaffold in situ.
 27. The composition of claim 26, wherein said composite composition further comprises: (i) a chemokine; (ii) a programming factor; (iii) a tumor antigen; (iv) a bioactive factor selected from the group consisting of vascular endothelial growth factor (VEGF), acidic fibroblast growth factor (aFGF), basic fibroblast growth factor (bFGF), placenta growth factor (PIGF), platelet derived growth factor (PDGF), leptin, hematopoietic growth factor (HGF), VEGF receptor-1 (VEGFR-1), VEGFR-2, a member of the bone morphogenetic protein (BMP) family, granulocyte/macrophage colony stimulating factor (GM-CSF), FMS-like tyrosine kinase 3 ligand (FIt3 ligand), hepatocyte growth factor, stromal derived factor 1 (SDF-1), insulin like growth factor (IGF), anti-VEGF antibody, anti-aFGF antibody, anti-bFGF antibody, anti-PIGF antibody, anti-leptin antibody, anti-HGF antibody, anti-VEGFR-1 antibody, antiVEGFR-2 antibody, anti-PDGF antibody, anti-BMP antibody, anti-FIt3 ligand, and anti-IGF antibody; (v) a bioactive factor selected from the group consisting of BMP-2, BMP-4, and RunX; (vi) a bioactive factor comprising MyoD; and/or (vii) a bioactive factor comprising FGF.
 28. The composition of claim 27, wherein: (i) said chemokine comprises granulocyte/macrophage colony stimulating factor (GM-CSF); and/or (ii) said programming factor comprises a condensed oligonucleotide, wherein said condensed oligonucleotide comprises CpG or plasmid DNA.
 29. The composition of claim 26, wherein said cells migrate into macropores of said composite composition in situ, wherein said cells comprise lymphocytes or antigen presenting cells, wherein said antigen presenting cells comprise dendritic cells.
 30. The composition of claim 26, wherein: (i) said sacrificial porogen hydrogel micro-beads comprise oxidized alginate; (ii) said sacrificial porogen hydrogel micro-beads comprise 3-7.5% oxidized alginate; (iii) said sacrificial porogen hydrogel micro-beads comprise alginate dialdehyde; or (iv) said sacrificial porogen hydrogel micro-beads comprise 20 mg/mL oxidized alginate and 7.5 mg/mL unmodified alginate.
 31. The composition of claim 26, wherein said sacrificial porogen hydrogel micro-beads or said bulk hydrogel comprise an isolated cell, optionally wherein said isolated cell is a mesenchymal stem cell, a myoblast, a vascular progenitor cell, a differentiated cell derived from an embryonic stem cell or an induced pluripotent stem cell, an induced pluripotent cell, or a cell that was directly reprogrammed from a fibroblast to a differentiated state.
 32. The composition of claim 26, wherein: (i) said sacrificial porogen hydrogel micro-beads comprise an elastic modulus of between 20 kPa and 60 kPa; (ii) said bulk hydrogel comprises a peptide comprising an amino acid sequence of PHSRN (SEQ ID NO: 1), DGEA (SEQ ID NO: 2), or RGD; (iii) said bulk hydrogel comprises a density of RGD peptides from 2 to 10 peptides per alginate polymer chain; and/or (iv) said bulk hydrogel comprises an initial elastic modulus of at least 40 kPa.
 33. The composition of claim 26, wherein: (i) said composite scaffold composition promotes bone or cartilage repair, regeneration, or formation; (ii) said composite scaffold composition further comprises a bioactive factor selected from the group consisting of BMP-2, BMP-4, and RunX; (iii) said sacrificial porogen hydrogel micro-beads or said bulk hydrogel comprise an isolated bone cell selected from the group consisting of an osteoblast, an osteocyte, an osteoclast, and an osteoprogenitor, optionally, wherein said isolated bone cell is an autologous or allogenic cell; and/or (iv) said sacrificial porogen hydrogel micro-beads or said bulk hydrogel comprise an isolated cartilage cell, optionally, wherein said isolated cartilage cell comprises a chondroblast.
 34. The composition of claim 26, wherein: (i) said composite scaffold composition promotes muscle repair, regeneration, or formation; (ii) said composite scaffold composition further comprises a bioactive factor, wherein said bioactive factor comprises MyoD; and/or (iii) said sacrificial porogen hydrogel micro-beads or said bulk hydrogel comprise an isolated muscle cell selected from the group consisting of a skeletal muscle cell, a cardiac muscle cell, a smooth muscle cell, and a myoprogenitor cell, optionally, wherein said isolated muscle cell is an autologous or allogenic cell.
 35. The composition of claim 26, wherein: (i) said composite scaffold composition promotes skin repair, regeneration, or formation; (ii) said composite scaffold composition further comprises a bioactive factor, optionally, wherein said bioactive factor comprises FGF; and/or (iii) said sacrificial porogen hydrogel micro-beads or said bulk hydrogel comprise an isolated skin cell selected from the group consisting of a fibroblast, a dermal cell, an epidermal cell, and a dermal progenitor cell, optionally, wherein said isolated skin cell is an autologous cell or an allogeneic cell.
 36. The composition of claim 26, wherein sacrificial porogen hydrogel micro-beads are present at a density of 60% of the overall volume of the composite scaffold composition.
 37. The composition of claim 30, wherein: (i) at least 5% of said alginate is oxidized; and/or (ii) said bulk hydrogel comprises unmodified alginate.
 38. The composition of claim 26, wherein: (i) said sacrificial porogen hydrogel micro-beads comprise an oxidized alginate polymer having a molecular weight from 5,000 to 500,000 Daltons (Da); (ii) said bulk hydrogel comprises an alginate polysaccharide having a molecular weight from 5,000 to 500,000 Da; (iii) said sacrificial porogen hydrogel microbeads comprise polymers with a molecular mass of approximately 50 kDa; (iv) said bulk hydrogel comprises polymers with a molecular mass of approximately 250 kDa. (v) dendritic cells are recruited into said macropores and programmed to be activated antigen-presenting dendritic cells to elicit an antitumor response; (vi) said sacrificial porogen hydrogel micro-beads and said bulk hydrogel are biodegradable. (vii) said macropores comprise macropores that are 50 μm to 500 μm in diameter; (viii) said macropores comprise macropores that are 100 μm to 500 μm in diameter; and/or (ix) said macropores comprise macropores that are 50 μm to 400 μm in diameter.
 39. The composition of claim 26, wherein: (i) said sacrificial porogen hydrogel micro-beads comprise oxidized alginate and said bulk hydrogel comprises oxidized alginate; (ii) said bulk hydrogel comprises less oxidized alginate than said sacrificial porogen hydrogel micro-beads; (iii) said sacrificial porogen hydrogel micro-beads comprise 3-7.5% oxidized alginate; and/or (iv) said sacrificial porogen hydrogel micro-beads comprise a shorter alginate polymer than said bulk hydrogel.
 40. A method of deploying cells from a scaffold into tissues of a mammalian subject, comprising administering to a subject a composition of claim 26, and wherein said sacrificial porogen hydrogel micro-beads or said bulk hydrogel comprise an isolated cell.
 41. A method of recruiting cells into a scaffold in vivo, comprising administering to a subject a composition of claim
 26. 